1. Field of the Invention
This invention relates to a radio frequency coil for use in a magnetic resonance imaging system. Such radio frequency coils typically operate at a frequency (or band of frequencies) within a range of between 3 and 64 MHz to couple radio frequency energy to and/or from body tissue or other objects located within an imaging volume of a magnetic resonance imaging system.
2. Related Art
The art of magnetic resonance imaging (MRI) is now well developed and several different types of MRI systems are commercially available. In all of them, some means is provided to produce a very strong static magnetic field H.sub.0 and controlled spatial magnetic gradients therein (e.g., along three mutually-orthogonal coordinate axes). The static magnetic field is typically of an approximately-homogenous nature within a predefined imaging volume and the controlled magnetic gradients are typically approximately linear with respect to spatial displacements therewithin.
A programmed sequence of radio frequency pulses is transmitted into body portions located within the imaging volume at predetermined frequencies or frequency distributions (typically all situated within the range of 3-64 MHz depending on the strength of H.sub.0) so as to selectively nutate the magnetic moment of certain nuclei by predetermined amounts in accordance with well-known nuclear magnetic resonance (NMR) principles. After cessation of such transmitted RF pulses, the NMR nutated atoms tend to relax back toward alignment with the static magnetic field H.sub.0 and, in the process, produce characteristic NMR RF signals. Such RF signals are received, detected and processed to thereafter produce a desired MRI image of the body portion located within the imaging area in accordance with any one of many known MRI techniques as will be appreciated by those in the art. The transmitted RF pulses typically are synchronized with a special sequence of current pulses passed through various magnetic gradient coils during the imaging process so as to effect spatial information encoding processes and/or to provide known types of NMR phasing control.
In some MRI apparatus, the static magnetic field H.sub.0 and/or the magnetic gradient coils are realized in the form of large solenoidal coils or, in the case of gradient coils, saddle-shaped coils conformed to a generally tubular configuration. In such cases, it is naturally necessary for patient access to the imaging volume to be provided only along a narrow tunnel through the tubular-shaped apparatus. With some patients, this may give rise to claustrophobic reactions. It also makes it extremely cumbersome to access the image volume (e.g., so as to adjust the relative positioning of RF transmit and/or receive coils or to attend to patient needs).
Other types of MRI systems utilize a pair of magnetic poles (e.g., permanent magnets or electromagnets with ferromagnetic poles and flux return paths) disposed on opposite sides of the image volume to create the requisite static magnetic field H.sub.0. Necessary magnetic circuits for return flux (i.e., outside the image volume) between the magnetic poles and/or the magnetic gradient coils (e.g., in a tubular form or flat) or decorative cover systems have been constructed in various ways. Early permanent magnet systems limited access to the image volume except along a generally tunnel-shaped area through which the patient was transported into the image volume. Thus, as with the solenoidal field generating devices, access to the image volume initially was essentially limited to only one or two aligned open and unobstructed patient access ports or areas--i.e., the opposite ends of the patient transport tunnel aligned with the patient transport axis.
The present applicant earlier discovered an improved magnetic resonance imaging apparatus wherein the static field magnet and gradient coil and decorative/functional outer cover structures are configured so as to leave an open and unobstructed patient access area communicating directly with the image volume along a direction perpendicular to the patient transport axis (e.g., see U.S. Pat. No. 4,829,252 issued May 9, 1989, the entire contents of which is hereby incorporated by reference). In the preferred exemplary '252 patent embodiment, such transverse access to the imaging volume may be had from two opposite sides of the patient transport mechanism while in yet another exemplary embodiment, such transverse access to the imaging volume passes through one side (or even the top) of the MRI system. In such exemplary embodiments, magnetic flux return circuits are preferably in the form of one or more cylindrical columns (e.g., four of them) disposed radially outwardly of the magnetic poles. In this manner, transverse unobstructed access to the imaging volume is provided not only along the patient transport axis, but also through at least one additional transverse port provided between such columnar return flux circuit structures. The new '252 patented system took unique advantage of an available open static magnet structure by coordinating gradient coil and housing structures so as to maintain such "openness" in the final completed MRI structure. That is, no obstructing housings or other structures were used to obstruct such transverse access path.
No matter how the strong magnetic field H.sub.0 is generated, such prior MRI systems have traditionally used essentially pure soft copper conductors for the RF coil structures. Such pure copper conductors are relatively "soft" in that they can be easily altered in shape if not otherwise supported to the desired shape for any given RF coil structure. For example, since the RF coil structures typically come into contact with or close proximity to the human body parts being imaged, it would be unrealistic to expect such soft copper conductors to retain their shape throughout an intended life cycle of repeated usage in practical commercial applications unless given an external supportive structure of some kind (e.g., an exterior supportive skeleton analogous that of exoskeletal members of the animal kingdom).
If RF coils were to be made of soft pure copper without any exterior support, the windings would be expected to move about in space (e.g., when coming into contact with normally expected forces during normal use). This would cause at least two different problems. First, the inductance of the RF antenna coil would change causing a variation in the coupling of a receiver preamp to the coil and therefore a change in the voltage seen by the preamp. Second, the change in coil size changes the emf induced in the rotating magnetic spins (or, by reciprocity the RF field made if the coil is being used as a transmitter). This effect also causes a variation in the received signal. If one were to have variation in signal from one MRI signal acquisition to the next, there would be an unacceptable bleeding of the imaged object in the phase-encoded direction.
Accordingly, it has heretofore been common practice to completely encase the RF coil conductors in an external, supporting, insulative body. For many MRI system environments, such relatively encumbered coil structures are less than optimum. For example, the solid insulating support for the coil conductor may serve to obstruct desired medical procedure access through the coil structure during set up, imaging or between imaging sequences of the MRI system. Furthermore, such a completely closed and solid appearance of the composite coil structure may also adversely enhance claustrophobic feelings or reactions of patients whose body parts may be encompassed by the thus relatively-closed coil structure.
Very stiff, potentially self-supporting electrical conductors are, of course, already known for other applications in the general field of electrical system design. For example, high-strength and high-conductivity copper-silver (Cu-Ag) alloys have been available for some time from Showa Electric Wire and Cable Company, Ltd. (Tokyo Toranomon Bldg., 1-1-18 Toranomon, Minato-ku, Tokyo, 105, Japan or Showa Electric America, Inc., Suite 1142, Russ Bldg., 235 Montgomery Street, San Francisco, Calif., 94104-3062). Such high-strength, high-conductive Cu-Ag alloys have approximately 80% the conductivity of pure copper--but with remarkably higher strength and stiffness. Other stiff (i.e., "hard") conductive materials which may be suitable include Beryllium-copper, phosphor-bronze and titanium. Even if the conductivity of some such stiff materials is less than Cu-Ag, they may still be in a useful range of conductivity. Furthermore, the conductivity may be increased by processing to avoid too much oxygen contamination. Other materials having relatively good conductivity high strength and without ferromagnetic content may also be presently available or may become available in the future.
Such high-conductive/high-strength conductors have been used previously for development of high-field magnets, such as pulse magnets. Such alloys apparently exhibit the combination of high strength and high conductivity due, at least in part, to proper combinations of cold working and heat treatments--but result in alloys that are easily melted and cast or forcefully bent into desired magnet coil structures for high-field magnets. Various standard sizes (e.g., bar lengths having cross-sections of 2.times.3 mm, 2.5.times.4 mm, 4.times.6 mm, etc.) are commercially available and are known to have excellent and uniform characteristics for use in high-field magnet applications.
Attempts apparently have been made previously to find other, more general purpose, applications and markets for this high-strength, high-conductivity Cu-Ag alloy conductive material. Either heat fusion or adhesive-type insulation can be supplied on the conductive bars as supplied by the vendor.
Although RF coils for MRI systems are often schematically depicted in drawings without external support so that one can visualize the individual internal conductors, in commercial practice, the actual RF coil structure utilized with the MRI system is necessarily substantially, if not completely, encased within an exterior supporting insulating body for reasons already noted above. Applicant's search of prior issued U.S. patents has discovered no prior attempt to actually achieve a self-supporting RF coil for an MRI system. For example, the following patents have been reviewed:
______________________________________ 4,636,729 Maurer et al (1987) 5,435,302 Lenkinski et al (1995) 4,620,155 Edelstein (1986) 4,649,348 Flugan (1987) 4,692,705 Hayes (1987) 5,235,283 Lehne et al (1993) 5,334,937 Peck et al (1994) 5,357,958 Kaufman (1994) 5,378,988 Pulyer (1995) 5,381,122 Laskaris et al (1995) 5,474,069 Wong et al (1995) 5,519,321 Hagen et al (1996) ______________________________________
Maurer et al direct their teaching to a "self-supporting" magnetic gradient coil system. However, while some of the gradient coil sub-structures appear to be truly "self-supporting" in the sense of (presumably insulated) conductors being directly cemented to one another, the entire gradient coil assemblage is only made "self-supporting" by the use of several longitudinal support elements of non-magnetic insulating material. The primary purpose of the open-type gradient coil structure is apparently an attempt to reduce audible noise within the MRI system during operation. In any event, there is nothing in this reference that would suggest the desirability of making RF coils in any way wholly or even partially "self-supporting".
Lenkinski et al (1995) is actually directed to a flexible RF surface coil--but contrasts such flexible structure against prior art structures said to be "rigid"--citing as examples Edelstein, Flugan and Hayes. However, the RF coil structures therein described are only "rigid" or "stiff", if at all, because they are formed on and attached to the outer surface of a rigid or stiff coil former (or encased in a rigid or stiff casement or the like). None of these references teach or suggest use of a rigid or self-supporting RF coil conductor.
Kaufman, Pulyer and Hagen et al are examples of MRI systems and/or components designed so as to retain some degree of patient access during the imaging process. Peck et al is an example of typical rigid coil former constructions used in MRI systems (this one with gradient coils). Collectively, these plus Lehne et al, Laskaris et al and Wong merely confirm that prior art MRI RF coils are invariably formed from conventional (soft) copper--and most typically made rigid or stiff only by associating the soft copper conductors with rigid external coil former structures.
There may be many reasons that those in the prior art have so far failed to teach or suggest the use of self-supporting stiff conductors for RF coils in an MRI system. For example, the Cu-Ag alloy conductors available from Showa are so very stiff that they are very difficult to bend by hand. Furthermore, since the conductivity of such material is necessarily less than pure copper, use of such material would be contrary to normal inclination for this reason as well. The possible presence of ferromagnetic impurities (e.g., iron, cobalt, nickel, etc.) in alloy conductors may be another reason that those in the prior art have not previously attempted use of such materials in an RF MRI coil application.